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1
ERODIBLE DRUG DELIVERY SYSTEMS FOR TIME-CONTROLLED 1
RELEASE INTO THE GASTROINTESTINAL TRACT 2
3
Alessandra Maroni, Lucia Zema, Matteo Cerea, Anastasia Foppoli, Luca Palugan, 4
Andrea Gazzaniga* 5
6
Università degli Studi di Milano, Dipartimento di Scienze Farmaceutiche, Sezione di 7
Tecnologia e Legislazione Farmaceutiche "Maria Edvige Sangalli", Via G. Colombo 71, 8
20133 Milan, Italy. 9
*Corresponding Author: [email protected], phone +39 02 503 24654 10
11
12
*Manuscript
2
Abstract 13
In oral delivery, lag phases of programmable duration that precede drug release may 14
be advantageous in a number of instances, e.g. to meet chronotherapeutic needs or 15
pursue colonic delivery. Systems that give rise to characteristic lag phases in their 16
release profiles, i.e. intended for time-controlled release, are generally composed of a 17
drug-containing core and a functional polymeric barrier. According to the nature of the 18
polymer, the latter may delay the onset of drug release by acting as a rupturable, 19
permeable or erodible boundary layer. Erodible systems are mostly based on water 20
swellable polymers, such as hydrophilic cellulose ethers, and the release of the 21
incorporated drug is deferred through the progressive hydration and erosion of the 22
polymeric barrier upon contact with aqueous fluids. The extent of delay depends on the 23
employed polymer, particularly on its viscosity grade, and on the thickness of the layer 24
applied. The manufacturing technique may also have an impact on the performance of 25
such systems. Double-compression and spray-coating have mainly been used, resulting 26
in differing technical issues and release outcomes. In this article, an update on delivery 27
systems based on erodible polymer barriers (coatings, shells) for time-controlled release 28
is presented. 29
30
Keywords 31
Oral pulsatile release, Oral colon delivery, Coating, Swellable/erodible hydrophilic 32
polymers, Injection-molding, Fused deposition modeling 3D printing. 33
34
3
Introduction 35
Oral delivery systems for time-controlled release are able to defer the onset of drug 36
release into the gastrointestinal tract for a programmable lag period independent of pH, 37
ionic strength, enzyme concentration and other physiological parameters. It is by now 38
recognized that a delay prior to release may be advantageous for effective 39
pharmacological treatment of several pathologic conditions [1]. This is typically the 40
case with a variety of high-morbidity rheumatic, cardiovascular and respiratory chronic 41
diseases, which show cyclic patterns in their signs and symptoms [1,2]. When these 42
mainly recur at night or in the early morning hours, bedtime administration of drug 43
products having a proper lag phase in their release profile would help provide 44
pharmacological protection as needed. On the other hand, both untimely awakenings, as 45
an immediate-release dosage form would require, and exposure to unnecessarily 46
sustained therapeutic drug levels, as prolonged-release formulations taken before sleep 47
would entail, could thereby be overcome. As a result, not only the efficacy and safety of 48
a treatment but also the relevant patient compliance may greatly be enhanced through 49
the use of chronopharmaceutical delivery systems. 50
Besides, a lag phase prior to release allows to target the colonic region with drug 51
molecules intended for either a local action, e.g. to treat Inflammatory Bowel Disease 52
(IBD), or for systemic absorption, especially of biotech molecules that pose stability 53
issues in the proximal gut and may benefit from the aid of enhancers for mucosal 54
permeation [3,4]. When colon delivery is sought, the lag phase is expected to last 55
throughout the entire small intestinal transit (3 h ± 1 SD), which was reported not to be 56
strongly influenced by the characteristics of dosage units and by food intake [5,6]. 57
Moreover, the lag period should be started upon emptying from the stomach rather than 58
4
on administration, owing to the high variability of gastric residence that cannot reliably 59
be predicted. Hence, in order to attain colonic release based on a time-controlled 60
approach, enteric coating is generally required. 61
Repeated lag phases, each followed by the release of a drug dose fraction, may be 62
exploited to fulfill multiple daily administrations regimens when prolonged release is 63
not a viable option, e.g. because of pharmacokinetic (strong first-pass effect) or 64
pharmacodynamic (tolerance) constraints. Successive release pulses are also proposed 65
as an alternative strategy in antibiotic therapy, possibly resulting in restrained growth of 66
resistant bacterial strains [7]. 67
Finally, properly modulated lag phases prior to the release of co-administered 68
bioactive compounds may avoid undesired drug-drug interactions in the gastrointestinal 69
tract and overcome the need for differing dosing schedules, thus improving the overall 70
patient convenience and compliance [8]. 71
Peroral delivery systems for time-controlled release are expected to yield lag phases 72
on the order of few hours, which may be consistent with their mean residence time 73
within the digestive tract. These are often pursued through functional polymeric barriers 74
that enclose an inner drug formulation [9,10]. According to the physico-chemical 75
properties of their polymeric components and type of excipients added (plasticizers, 76
pore formers, bulking agents), such barriers delay the onset of release via differing 77
mechanisms. They may indeed undergo time-programmed disruption, become leaky or 78
be subject to progressive erosion/dissolution. In particular, erodible systems are 79
generally single-unit dosage forms based on a drug containing-core, such as an 80
immediate-release tablet or capsule, and a swellable hydrophilic barrier of adequate 81
thickness and polymer viscosity. Such a barrier may be a coating or, in more recent and 82
5
innovative instances, a freestanding release-modifying shell available for filling with 83
any drug formulation. 84
Because of the inherent safety and biocompatibility profile as well as of their 85
availability in a range of grades and reasonable costs, hydrophilic cellulose derivatives, 86
namely hydroxypropyl methylcellulose (HPMC) and, less frequently, hydroxypropyl 87
cellulose (HPC) and hydroxyethyl cellulose (HEC), are broadly used as the functional 88
polymers in erodible delivery systems [11]. Other polysaccharides, including 89
galactomannans, alginates, xanthan gum, and non-saccharide hydrophilic polymers, 90
such as polyvinyl alcohol (PVA) and polyethylene oxide (PEO), are nonetheless also 91
employed. All of these materials are largely utilized in the food, pharmaceutical, 92
nutraceutical and cosmetic industries mainly as rheology-modifiers, stabilizers, binders 93
and film-coating agents. 94
Upon water uptake, such polymers typically go through a glassy-rubbery 95
thermodynamic transition that is associated with distension and disentanglement of their 96
macromolecular chains [12-14]. Consequently, the polymer structure may expand, erode 97
due to mechanical attrition and/or dissolve at a rate that chiefly depends on the relevant 98
physico-chemical characteristics and on the ionic strength and temperature of the 99
medium. As the aqueous fluid penetrates into the polymeric layer, a swelling front, i.e. 100
the boundary between the glassy and the rubbery domain, and an erosion front, at the 101
interface between the rubbery polymer and the outer medium, are identified. Depending 102
on the relative movements of the swelling and erosion fronts, which in turn are 103
governed by the hydration, dissolution and viscosity properties of the polymer, a gel 104
layer of varying thickness is formed. 105
6
In a few instances, insoluble materials are added to the hydrophilic polymers to 106
modulate the degree of hydration of the barrier, or even used as the main components of 107
mechanically erodible coatings. In the latter case, their erosion in aqueous fluids would 108
need to be promoted by surfactant excipients. 109
Drug release from hydrophilic erodible systems is in principle deferred until the 110
entire polymeric layer is in the swollen state, i.e. when the swelling front has reached 111
the drug core, possibly followed by extensive dissolution/erosion of the hydrated 112
polymer. The duration of the lag phase is indeed dictated by the physico-chemical 113
properties of the polymer employed, primarily molecular weight and degree of 114
hydrophilicity, and by the thickness of the erodible barrier. The manufacturing 115
technique, which may range from double-compression and spray-coating to hot-116
processing, can also affect the layer functionality. 117
In the following sections, oral delivery systems for time-controlled release provided 118
with an erodible polymer barrier are reviewed, and advances in this particular field are 119
illustrated with special emphasis on formulation and performance issues. 120
121
Erodible systems manufactured by double-compression 122
The manufacturing of oral delivery systems provided with erodible coatings dates 123
back to the early 90s. Until then, the use of such polymers in the manufacturing of solid 124
dosage forms was tied to tableted hydrophilic matrices for prolonged release. Indeed, 125
double-compression technique, also known as press-coating, was adopted in all initial 126
attempts. The first one concerned a three-layer tablet system that was proposed for two-127
pulse release of drugs [15,16]. Such a system was composed of two conventional drug 128
(ibuprofen) layers and a high-viscosity HPMC (Methocel® K4M and Methocel
® K15M) 129
7
barrier in between. An impermeable ethylcellulose (EC) film covered the lateral area 130
and one of the bases of the assembly so that the outer surface of a single drug layer was 131
allowed to interact with solvent upon first contact with the medium. The former dose 132
fraction was thereby released, whereas the latter was released after a lag phase due to 133
the hydration and erosion of the polymer barrier. The delay between the release pulses 134
depended on the viscosity of the polymers employed, and release of the latter dose 135
fraction was slower. This was ascribed to a less efficient activation of the disintegrant 136
incorporated within the inner drug layer that was progressively exposed to the aqueous 137
fluid. The release behavior observed in vitro was reflected in two-peak plasma 138
concentration curves in healthy volunteers. However, because of its multiple-layer 139
configuration and the need for a partial coating, the system would involve serious 140
scalability issues. Therefore, a simpler press-coated formulation was designed, wherein 141
the polymer, a low-viscosity HPMC (Methocel® K100 LV), covered the entire surface 142
of the core [17]. The coated system could yield single-pulse release after a lag phase or, 143
administered in combination with an immediate-release tablet, the repeated release 144
performance attained from the previous device. In the double compression process, 145
positioning of the core tablet in the die represented a critical step. However, by correctly 146
centering it within the polymer powder bed, biconvex tablets with coatings of 147
homogeneous thickness were obtained. As desired, the in vitro release was delayed for a 148
reproducible period of time, although leaching of a small percentage of the drug content 149
prior to the quantitative release phase was inferred from the curves. This was ascribed to 150
premature outward diffusion of dissolved drug molecules through the swollen polymer 151
coating. 152
8
A low- and a high-viscosity HPMC grade (Methocel® K100 and Methocel
® K4M) 153
were used, either alone or mixed with each other, as the coating agents of a delivery 154
system containing ibuprofen, aimed at the chronotherapy of rheumatoid arthritis, or 155
pseudoephedrine hydrochloride, a water-soluble model drug [18-20]. Increasing the 156
coating level or the amount of high- vs low-viscosity polymer resulted in longer lag 157
times and slower in vitro release as well as decreased absorption rates in healthy 158
volunteers. Sodium alginate, as compared with HPMC, performed as a less effective 159
barrier-forming polymer. Incorporation of a fraction of the drug dose in the coating 160
layer changed the release behavior, generally yielding biphasic kinetics that depended 161
on the composition of the polymeric coat and its drug load. 162
High-viscosity HPMC (Methocel® K4M, Methocel
® K15M and Methocel
® K100M) 163
was employed to prepare a system intended for colonic delivery of the anti-parasitic 164
drug tinidazole [21]. An enteric coating was applied externally to enable site-selective 165
release. The lag phase duration and the release rate were markedly affected by the 166
viscosity grade of the polymer, while hardness of press-coated tablets in a 40-60 N 167
range did not impact on the relevant performance. Administered to 2 healthy volunteers, 168
the system was shown to disintegrate in the ascending colon. Methocel® K100M was 169
also used to coat, at a compression force of 60-80 N, minitablet cores (3 mm in 170
diameter) intended for immediate or prolonged release of nifedipine [22]. By combining 171
differing core and coated formulations in a gelatin capsule, a variety of release patterns 172
were achieved. 173
Low-viscosity HPMC coatings were applied by an alternative tableting method 174
(One-Step Dry-Coated, OSDRC) based on the use of a specially modified equipment, 175
which was previously set up in order to overcome disadvantages typically encountered 176
9
with conventional double-compression technique [23]. These mainly encompass the 177
need for poorly scalable multiple-step processing, the issue of coat thickness 178
homogeneity and difficulties in attaining relatively low coating levels. By the OSDRC 179
method, layers of 0.5-2 mm were obtained, with satisfactory thickness homogeneity and 180
practically unchanged performance within a 100-200 MPa range of compression 181
pressure. 182
High-viscosity HPMC was mixed with polyvinylpyrrolidone (PVP) at different ratios 183
and applied, by conventional double-compression technique, to minitablet cores 184
containing solid felodipine/PVP dispersions [24,25]. Mixing with PVP at 30-50% 185
resulted in improved mucoadhesion of the HPMC coating. The delays prior to a rapid 186
release of the drug increased in duration with the percentage of HPMC in the 187
formulation. In vitro delays of more than 10 h were observed with amounts of HPMC at 188
which mucoadhesive properties were enhanced. The issue of possible inconsistency 189
between duration of the lag phase and gastrointestinal transit was faced by the design of 190
a floating pulsatile delivery system aimed at gastro-retention [26]. For this purpose, a 191
verapamil hydrochloride tablet was first coated with low-viscosity HPMC (Methocel® 192
E5, Methocel® E15 or Methocel
® E50), expected to defer the onset of drug release. A 193
blend of a high-viscosity grade of the polymer (Methocel® K4M) and Carbopol
® 934P, 194
which also contained sodium bicarbonate to generate effervescence, was subsequently 195
applied to a single face of the unit coated with low-viscosity HPMC. The system was 196
proved able both to delay the onset of release and to float in vitro. Lag time depended 197
on the viscosity and amount of HPMC in the coating. A ɣ-scintigraphic evaluation in 6 198
healthy volunteers highlighted the extended gastric residence of the dosage form and 199
reproducible lag phases before release. In all cases, this occurred in the stomach or 200
10
small intestine. Recently, various grades of HPMC were used to coat tablet cores based 201
on drugs with differing solubility values [27]. Poorly soluble carbamazepine was 202
released in a pulsatile fashion after erosion of the coating polymer, and the viscosity 203
characteristics of the latter strongly impacted on the relevant performance. On the other 204
hand, more soluble drugs were released in a sigmoidal mode, which was attributed to 205
their diffusion through the fully hydrated HPMC layer, and a poor influence of the 206
polymer viscosity was noticed. The outward diffusion of the drug prior to its 207
quantitative release could be prevented by inserting an enteric film below the erodible 208
coating. However, this would ultimately impart pH-dependence to the lag phase and 209
possibly hamper a timely release of the drug for chronotherapeutic purposes. The 210
amount of HPMC also affected the time and rate of release. 211
Although HPMC was most widely utilized as a coating agent intended for delaying 212
dug release, the use of other hydrophilic cellulose derivatives was reported. Particularly, 213
HPC was the component of a compressed shell that was separately prepared and, once 214
perforated, manually assembled with a cylindrical core tablet containing isosorbide-5-215
nitrate [28]. The upper and lower bases of the resulting system were coated with an 216
impermeable ethylene vinyl acetate copolymer film. Release was deferred until the 217
polymeric shell was completely eroded or detached. Lag time was affected by the 218
thickness of such a shell and by the composition of the core. Indeed, replacing 219
microcrystalline cellulose with lactose shortened the lag phase because of the osmotic 220
effect exerted by the latter filler. 221
A diltiazem hydrochloride system based on HPC was prepared by conventional 222
press-coating [29,30]. As with HPMC, the lag phase duration was modulated either by 223
increasing/decreasing the amount of coating material applied or by employing HPC, or 224
11
mixtures thereof, with differing viscosity values. Prototype formulations having in vitro 225
delays of approximately 3 h and 6 h were administered to beagle dogs. A good 226
agreement was found between in vitro and in vivo data relevant to the former prototype, 227
whereas lag time in vivo was shorter than in vitro in the latter case. This gap was 228
reduced when a paddle rotation speed of 150 rpm was set instead of 100 rpm during 229
release testing. In order to assess its potential for colon delivery, the system having lag 230
time of 3 h was provided with an enteric HPMCAS film containing a gastric emptying 231
marker (phenylpropanolamine hydrochloride) [30]. The mean difference between the 232
time of first appearance in plasma (TFA) of the drug and of the marker molecule was of 233
about 3 h, which was consistent with the lag time obtained from the pH 6.8 fluid stage 234
of the in vitro test. HPC was also used in admixture with EC at a weight ratio of 7:1 235
[31]. The addition of the insoluble polymer aided a faster release of aceclofenac, 236
intended for the chronotherapy of rheumatic morning pain, after the delay period. The in 237
vitro performance of press-coated systems based on this blend was proved independent 238
of various parameters, such as the compression force, paddle rotation speed during 239
release testing and pH of the medium. Provided with an enteric-coating, the formulation 240
was administered to rabbits, showing a clear lag phase as opposed to an immediate-241
release tablet. However, due to variable residence of solid dosage forms in the stomach, 242
gastroresistance may prevent the anti-inflammatory drug from being released at the time 243
the disease symptoms occur. 244
Low-substituted HPC (L-HPC), an insoluble swellable hydrophilic cellulose ether 245
that is largely used as a disintegrant, was mixed with glyceryl behenate at differing 246
ratios and subjected to a melt-granulation process [32]. The resulting granules were 247
applied by double-compression to theophylline tablet cores to give the erodible layer. 248
12
The press-coated tablet was studied in vitro and in beagle dogs, by pharmacokinetic as 249
well as ɣ-scintigraphic techniques. Lag phases were reproducible in duration and 250
increased with the amount of glyceryl behenate in the coating formula up to 75%. No 251
significant differences were found either between in vitro and in vivo lag times, or 252
between in vivo lag and disintegration times, both in the fasted and fed state. 253
Press-coated tablets for chronotherapeutic purposes were prepared from HEC 254
employed as the erodible barrier-forming material [33]. The onset of release of 255
diltiazem hydrochloride from the core was delayed in vitro as a function of the coating 256
level and the viscosity grade of HEC. The particle size of the polymer also affected lag 257
time. Using powders with larger particle dimensions was associated with shorter delay 258
phases, which was ascribed to the positive effect of a greater porosity on the polymer 259
hydration process. The role played by HEC viscosity was studied in healthy volunteers 260
[34]. When this parameter increased, progressively longer lag time (Tlag) and lower 261
maximum concentration (Cmax) values were observed in the plasma concentration vs 262
time curves. However, the area under the curve (AUC0-24 h) did not change significantly. 263
In vitro and in vivo lag times were in agreement. 264
Besides cellulose derivatives, the use of PEO as a hydrophilic erodible coating agent 265
was reported. Blended with PEG 6000 at 1:1, it was applied to tablets containing 266
acetaminophen and differing water soluble excipients, such as PEG 6000, sucrose and 267
lactose [35]. These were added in order to promote erosion of the core in the distal 268
intestine, where the press-coated tablets would be intended to release their drug load, 269
thus possibly counterbalancing the paucity of water of regional fluids. The core erosion 270
was experimentally quantified and expressed by a purposely introduced parameter, i.e. 271
the core erosion ratio. In a pharmacokinetic study conducted with fasted beagle dogs, 272
13
greater Cmax and AUC values were obtained from formulations having a higher core 273
erosion ratio. The amount of PEG 600 vs PEO was raised up to 5:1 in the coating of 274
nifedipine tablets containing sucrose as an erosion enhancer [36]. In vitro lag times 275
increased with the percentage of PEO and were aligned with TFA data in beagle dogs. 276
PEO formed the swelling/erodible upper layer of a press-coated system with an 277
impermeable cellulose acetate propionate shell covering one of the bases and the lateral 278
surface [37]. The amount of polymer in the top coating affected both the time and rate 279
of release of drug molecules with different solubility. Visual monitoring of 280
morphological changes undergone by the system during in vitro testing highlighted 281
gradual expansion and erosion of the partial PEO coat until final detachment from the 282
underlying unit. Used in place of PEO, sodium alginate and sodium 283
carboxymethylcellulose had less and greater impact on the release performance, 284
respectively, consistent with their observed swelling/erosion behavior. Guar gum having 285
ten-fold higher viscosity than PEO also exerted a tighter control of the onset and rate of 286
release [38]. Increasing the core diameter or adding a soluble filler, such as lactose, to 287
PEO or guar gum top layers resulted in reduced duration of the lag phase and enhanced 288
release rate. Differing PEO grades were employed to coat tablets containing solid 289
dispersions of indomethacin in a novel sucrose fatty acid ester carrier [39]. In vitro lag 290
time depended on the viscosity and amount of the coating polymer. In 6 healthy 291
volunteers, press-coated tablet systems with in vitro lag phase of approximately 6 h 292
brought about delayed appearance of indomethacin in plasma with respect to an 293
immediate-release commercial product. However, no significant differences were found 294
in the Cmax and AUC relevant to the two formulations. 295
14
Hydrophilic polymers of natural origin were also proposed as press-coating agents 296
for time-controlled delivery systems. For instance, powders composed of sodium 297
alginate and chitosan, forming a polyelectrolyte complex, and of lactose as a filler were 298
obtained by spray-drying, evaluated for flowability and compaction properties and 299
finally applied to acetaminophen tablets [40]. Through progressive erosion of the 300
coating layer, drug release was delayed in pH 6.8 fluid for a time interval that depended 301
on the chitosan content of the composite powder and on the polymer degree of 302
deacetylation. A prompt release phase was eventually observed. In pH 1.2 fluid, 303
acetaminophen was released slowly after longer delays. Prepared for comparison 304
purposes, physical mixtures of chitosan with spray-dried alginate/chitosan particles and 305
spray-dried powders composed of lactose and of pre-formed alginate/chitosan complex 306
failed to provide the desired release pattern. 307
Blends of the bacterial exopolysaccharide xanthan gum and plant galactomannan 308
locust bean gum were used in the double-compression coating of the SyncroDose™ 309
delivery system according to TIMERx® technology [41]. Differing release modes and 310
lag times were achieved by modifying the concentration and ratio of the two 311
polysaccharides, performing as synergistically interacting heterodisperse polymers. 312
313
Erodible delivery systems manufactured by spray-coating 314
The feasibility of coating techniques other than double-compression was explored for 315
the manufacturing of erodible polymer barriers able to control the onset of drug release. 316
Particularly, the goals were to establish simpler processing modes, with better industrial 317
scale-up prospects, exploit conventional production equipment and broaden the range of 318
viable core formulations (e.g. large tablets, minitablets, granules, pellets, gelatin 319
15
capsules) [17]. Furthermore, some performance issues, strictly connected with the 320
structure of press-coatings, their relatively high thickness and the relevant homogeneity 321
limitations, needed to be improved. These primarily involved extended, variable and 322
poorly flexible lag times, incomplete suppression of drug leakage during the delay 323
period and impact on the subsequent release phase. Preliminary spray-coating trials 324
were thus undertaken because such a technique would have allowed continuous and 325
uniform films to be formed rather than layers of pressed powder, and fluid bed as well 326
as rotating pan equipment to be utilized instead of specially devised or modified 327
tableting machines [17,42,43]. In addition, it could in principle be adapted to substrate 328
dosage forms having diverse size, surface and density characteristics, thereby 329
circumventing the dimensional and mechanical constraints associated with double-330
compression. A limited technical background was available on the use of swellable 331
hydrophilic polymers as film-coating agents, and this mainly concerned application of 332
low-viscosity grades as thin layers with protective, taste-masking or cosmetic function. 333
Nonetheless, HPMC with marked viscosity (Methocel® K4M and Methocel
® K15M) 334
appeared potentially suitable for delaying drug release for a time interval on the order of 335
hours without binding to excessively thick coatings. The polymers were suspended in a 336
hydro-alcoholic vehicle in order to counteract the thickening effect they exert upon 337
hydration. The ratio between ethanol and water needed to be adjusted so as to enable 338
nebulization of the coating suspension at reasonable rates and polymer concentrations 339
on the one hand, and adequate coalescence of the solid particles on solvent evaporation 340
on the other. The addition of plasticizing, anti-tacking and binding excipients, such as 341
PEG 400, talc and PVP, was investigated. The coatings applied to tablets and 342
minitablets were provided with consistent thickness and smooth surface. Moreover, they 343
16
yielded the desired release pattern. Considering the regulatory issues raised by organic 344
solvents, the feasibility of aqueous spray-coating by fluid bed was then evaluated using 345
HPMC having increasing viscosity, namely Methocel® E5, Methocel
® E50 and 346
Methocel® K4M [44-46]. The operating conditions, above all spray rate, inlet air 347
temperature and polymer concentration in the solutions, required an attentive set-up in 348
order to overcome major problems of powdering and nozzle clogging as well as lengthy 349
processing. The viscosity grade of the polymer chiefly affected the process time, 350
nebulization being possible only with diluted solutions that increased the spraying and 351
drying duration. From all of the polymers under investigation, coated units with 352
satisfactory physico-technological characteristics were obtained. The release behavior 353
was studied by paddle dissolution and modified disintegration apparatus. The latter 354
proved indeed better suited to prevent sticking of swollen HPMC to the vessels, thus 355
providing more reliable data. By both testing methods, a prompt release after a lag 356
phase was highlighted, which depended on the coating level and the polymer viscosity. 357
Using Methocel® E50 resulted in acceptable process feasibility, ability to delay drug 358
release and fine-tuning of the lag phase. Moreover, the coating process was shown 359
robust and potentially scalable. In the case of Methocel® K4M, not only the coating 360
operations were strongly impaired by the high viscosity of water solutions, but also a 361
small amount of drug was slowly released from coated units toward the end of the delay 362
period. This was attributed to the formation of a firm, poorly erodible gel structure 363
ultimately rupturing with the aid of the inner tablet disintegration upon water influx 364
[47]. When the Methocel® E50-based coating procedure was applied to hard- and soft-365
gelatin capsules instead of tablets, the process parameters needed to be adjusted in order 366
to prevent the sticking and shrinking of the shells [48]. In order to streamline 367
17
manufacturing of Methocel® E50-coated systems, alternative techniques, such as 368
tangential-spray film-coating and powder-layering carried out by fluid bed 369
rotogranulator, were attempted. Preliminary studies in volunteers demonstrated that, 370
irrespective of the core dosage form, delivery systems coated with Methocel® E50 by 371
aqueous spray-coating were able to defer drug appearance in saliva as a function of the 372
coating level [45,48]. The in vitro and in vivo lag phases were comparable in duration. 373
Moreover, when provided with a gastroresistant film, labeled formulations were shown 374
to consistently break up in the ascending colon. After low-molecular weight drugs, 375
chosen as models because of their stability characteristics and easy analysis, the 376
possibility of conveying bovine insulin by this delivery system was explored [48-52]. In 377
order to increase the chances of preserving integrity of the protein and promoting its 378
permeation through the intestinal mucosa, enzyme inhibitor and absorption enhancer 379
adjuvant compounds were incorporated in the formulation. Insulin was proved to 380
withstand all manufacturing steps, as inferred by assaying the degradation products 381
mentioned by European Pharmacopoeia, and was released in vitro in a pulsatile mode, 382
as previously observed with antipyrine and acetaminophen, along with the adjuvants. 383
The latter were also applied as a separate film enclosed between two Methocel® E50 384
layers, so that their release would occur earlier than that of the protein drug contained in 385
the core, and less threatening conditions could be established in vivo beforehand 386
[53,54]. 387
When erodible coatings were applied to minitablet cores, relatively larger amounts of 388
polymer were found necessary than with single units in order to obtain lag times 389
potentially suitable for chronotherapeutic or colonic release purposes [55-57]. Thus, the 390
thickness of the resulting film coatings would ultimately fail to comply with the size 391
18
requirements of multiple-unit dosage forms. Besides, depending on the viscosity grade 392
of the polymer employed, the rate of release at the end of the lag phase would most 393
likely be reduced. With the aim of overcoming this formulation issue, the external 394
application of an insoluble, flexible and increasingly leaky film was proposed. Such a 395
film was mainly intended to slow the uptake of water by the underlying HPMC layer, 396
and consequently the relevant hydration as well erosion processes, without acting as a 397
major mechanical constraint to the polymer expansion. Eudragit® NE 30 D was selected 398
as the film-forming agent, whereas various superdisintegrants, above all Explotab® V17, 399
were added as especially effective non-conventional pore formers. After tuning the 400
composition of the outer film and the ration between HPMC and polymethacrylate 401
coating levels, the desired release performance and dimensional characteristics were 402
obtained from formulations based on this novel two-layer design. 403
HPMC barriers derived from coalescence of polymer particles were prepared not 404
only by spray-coating but also by dipping, which circumvented the technical difficulties 405
associated with nebulization of highly viscous polymeric solutions [58]. Ethanol/water 406
mixtures were used to disperse the HPMC powder. Immersion steps, each followed by 407
manual hot-air drying, were repeated until the tablets had reached the established weight 408
gain. The latter was related to the lag phase duration. By affecting the structure of the 409
coat layer, parameters such as the ethanol/water volume ratio, the concentration of the 410
polymer and the time during which it was allowed to swell in the hydro-organic vehicle 411
also impacted on the release of nifedipine from the core tablet. 412
Waxy materials of natural origin, in admixture with a surfactant, were employed as 413
an alternative to swellable hydrophilic polymers in order to attain erodible barriers for 414
time-controlled release [59]. Spraying of water dispersions of such lipophilic coating 415
19
agents required that the processing temperature be set at relatively high values (75°C). 416
The resulting delivery system (Time Clock
) proved suitable for deferring salbutamol 417
sulfate release in vitro and in healthy volunteers. In both cases, the lag phases were 418
clearly dependent on the coating level. An agreement between in vitro and in vivo data 419
was achieved when media having increased viscosity were used for release testing, 420
which led to longer in vitro delays. The performance of the system in the 421
gastrointestinal tract was demonstrated not to be influenced by food intake in 6 subjects, 422
and AUC0-∞ as well as Cmax in the fasted state were consistent with those of an 423
immediate-release reference product. The Time Clock
system provided time-based 424
colonic delivery in humans when in gastro-resistant configuration, as highlighted by ɣ-425
scintigraphy [60]. This was confirmed in 8 fed volunteers through pharmaco-426
scintigraphic evaluation of a 5-aminosalicylic acid-containing formulation [61]. 427
428
Erodible delivery systems manufactured by hot-processing techniques 429
Hot-processing techniques, which enable the production of high-density structures of 430
any desired form from softened/melted thermoplastic material substrates, are raising 431
huge interest in every manufacturing area. However, their exploitation in the 432
pharmaceutical field is still fairly limited despite the enormous potential held [62-65]. It 433
is only recently that drug delivery applications mainly of hot-melt extrusion (HME), 434
injection-molding (IM) and three-dimensional (3D) printing by fused deposition 435
modeling (FDM) have been investigated and reported. Interestingly, the use of such 436
techniques was proposed for the production of void functional capsule shells 437
independent of their core units, with considerable prospective advantages from both the 438
technical and the regulatory point of views [66-68]. In this respect, the feasibility of IM 439
20
in fabrication of erodible shells intended to defer release of their contents was explored 440
[66]. HPC of various viscosity grades was selected as the capsule-forming polymer 441
because of the inherent thermoplastic behavior upon heating. A bench-top IM press was 442
employed, and the design of a specially suited mold was required. Through its use, cap 443
and body items were obtained within single automated production cycles. In vitro 444
studies pointed out a rapid release of the model drug after lag times that, composition 445
being equal, correlated with the thickness of shells in the 300-900 µm range 446
investigated. By visual inspection of capsule systems immersed in deionized water, it 447
was inferred that release after the delay phase would be connected with rupturing of the 448
hemispherical top and bottom ends of the device that were thinner than the cylindrical 449
region where cap/body portions overlapped. On administration of these prototypes to 3 450
healthy volunteers, the in vivo lag times calculated from salivary concentration curves 451
of acetaminophen were found in linear relationship with the in vitro ones [69]. The 452
design of a novel mold for 600 µm thick units and concomitant setting up of proper 453
formulation as well as operating parameters were subsequently undertaken [70]. This 454
allowed faster production cycles to be carried out without adding external or internal 455
lubricants. The shells obtained showed improved mechanical properties, which would 456
aid large-scale filling by the equipment used with conventional gelatin capsules, and 457
less variable thickness that was also closer to the theoretical value. Besides, the issue of 458
thicker body/cap overlap areas was overcome. As a result, more reproducible release 459
profiles were attained. The time to shell opening was demonstrated consistent 460
irrespective of differing types of solid dosage forms conveyed (fine powder, granules, 461
pellets, solid dispersion). These HPC capsules were successfully subjected to enteric 462
coating, with no need for sealing the assembled caps and bodies, and then to final curing 463
21
[71]. Such systems fulfilled the requirement of resistance in pH 1.2 medium for 2 h, 464
while maintaining the original pulsatile release curves when tested in pH 6.8 phosphate 465
buffer. Accordingly, they appeared potentially suitable for time-dependent colon 466
delivery, provided that the shell thickness be properly modulated so that duration of the 467
in vivo lag phase would match the small intestinal transit time. 468
Capsule shells composed of HPC were lately replicated by FDM 3D printing, starting 469
from filaments purposely prepared in-house by HME [68]. After assessing the 470
possibility of attaining hollow structures by the use of FDM and developing the needed 471
computer-aided design (CAD) files, bodies and caps of the shells were manufactured. 472
Overall, these exhibited satisfactory physico-technological characteristics and, 473
assembled into a drug-containing device, the typical lag phase before a rapid and 474
quantitative release. Upon contact with deionized water, the behavior of capsule shells 475
fabricated by FDM was comparable with that of analogous molded systems, thus 476
supporting the real-time prototyping potential of this 3D printing technique and its 477
possible exploitation in formulation development studies aimed at IM production. 478
Based on the expertise gained from the manufacturing of functional capsule shells, 479
cylindrical dosage forms, such as immediate- and prolonged-release polymeric units, 480
were also fabricated by HME and IM [72,73]. The relevant production via hot-481
processing was found to offer inherent advantages over the established techniques. 482
483
Conclusions 484
Drug delivery systems able to incorporate a lag phase of pre-established duration in 485
their release patterns are a topic of high current interest, primarily in connection with 486
oral chronotherapy and colon targeting. 487
22
Among the numerous formulation strategies proposed, those based on erodible 488
polymeric barriers have largely and successfully been exploited. As the main 489
components of such barriers, swellable/erodible polymers of hydrophilic nature, such as 490
HPMC and other cellulose derivatives, have especially been used. Indeed, they easily 491
enable fine-tuning of the release performance in terms of time and also rate through 492
proper selection of the type and amount of polymer, which will affect the thickness and 493
viscosity of the layer upon hydration. Erodible barriers intended for time-controlled 494
release generally consist in coating layers. These may partially or entirely enclose a 495
drug-containing core thus preventing it from immediately being exposed to aqueous 496
fluids on administration of the dosage form. Coatings may be applied by differing 497
techniques and, accordingly, possess diverse structural and functional characteristics. 498
Apart from coating layers, which are necessarily associated with a specific core 499
formulation, polymeric barriers in the form of erodible shells have recently been 500
manufactured by hot-processing, namely via IM and FDM. Because of the great 501
versatility in terms of design, high innovative content, excellent scale-up prospects and 502
unique benefits related to a separate development as well as production, these capsule 503
shells may open up new ways in the field of time-controlled release and, more broadly, 504
in the oral delivery area. 505
506
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